Method and assembly for magnetic resonance imaging and catheter sterring

ABSTRACT

An imaging and catheter steering assembly comprising a magnetic field generating assembly ( 1 - 6 ) operable in a first mode to generate a first magnetic field in a working volume ( 7 ) located outside the assembly, the first magnetic field being suitable for use in a catheter, steering procedure. In a second mode the assembly ( 1 - 6 ) generates a second, static magnetic field in the working volume ( 7 ) suitable for conducting a magnetic resonance imaging process (MRI), the second magnetic field being more uniform in the working volume than the first magnetic field. A catheter ( 67 ) has a magnetic seed attached whose orientation, and hence the steering direction of the catheter, is determined by interaction with the first magnetic field.

Nuclear Magnetic Resonance (NMR) and specifically Magnetic ResonanceImaging (MRI) is the established imaging method of choice for many typesof clinical diagnosis due to its exemplary soft tissue definition.Conventional whole body imaging systems generally use a superconductingsolenoid or Helmholtz-type coil pair (“C Magnet”) to generate therequired strong and uniform static magnetic field (called the B0 field).Patients undergoing examination lie within the bore of the solenoid orbetween the poles of the C-magnet. It is becoming increasingly desirableto monitor the progress and results of surgical procedures, such asbiopsy and intravascular catheterisation using MR imaging. This processis often called interventional MRI, or I-MRI. Surgeons have restrictedaccess to patients within conventional MRI apparatus, particularlysolenoid magnets; this has hampered the widespread application of I-MRItechniques.

Many surgical procedures previously requiring large scale opening of thebody cavity or brain-case are now routinely carried out using “keyholesurgery”. In this technique, often termed minimally invasive surgery(MIS), an endoscopic instrument or catheter is inserted through a smallincision and remotely operated by a highly skilled surgeon. A relatedarea is the new field of magnetically guided MIS, in which strongmagnetic fields are used to manipulate and guide surgical instrumentswithin the body. In particular, there have been recent advances inapparatus to remotely guide a catheter through a patient's vascularsystem by applying an external magnetic field with user-oriented fluxdirection. This field induces a torque on a permanent magnet “seed”implanted in the tip of a catheter, and thereby orients the tip in thedesired direction. When the tip is correctly aligned, forward motion isprovided by the surgeon pushing on the catheter from outside. In thisway a catheter inserted at the groin can be guided into the heart orbrain, obviating the need for traumatic opening of the chest orbrain-case. The external magnetic field is generated by a set of three(or more) orthogonal superconducting coils. An example of suitableapparatus is described in U.S. Pat. No. 6,241,671. Clearly thisapparatus limits direct access to the patient in the similar way toconventional MRI apparatus. (Note that the magnetic field generated bythe apparatus of U.S. Pat. No. 6,241,671 is dominated by large gradientsand is therefore entirely unsuitable for MR imaging).

A further related technology is monitoring the position of catheterswithin the patient's body using imaging. Commonly used techniques areX-ray fluoroscopy and ultrasound. X-ray fluoroscopy is particularlysuitable for real-time imaging, as the catheter material has quitedifferent X-ray absorption to tissue and is readily apparent in theimages. Monitoring catheter position using MRI is more difficult becausethe catheter generates no measurable MR signal (only NMR signals fromliquid sources are measured by conventional MRI hardware), and istherefore only visible by its contrast when immersed in tissuegenerating high NMR signal. Furthermore, whilst video-rate MR images arepossible, they demand high-specification hardware, so real time cathetertracking using MRI is difficult. However, MRI has several advantages fortissue imaging compared to X-ray (see below) and it is often onlynecessary to take “snap-shots” of catheter position at certain criticalstages of the surgical operation. Therefore many methods have beensuggested to monitor catheter positions using MRI. These include usingthe susceptibility artefact created by the catheter to make it visible(i.e.: detecting local distortion of the B0 field) (for example U.S.Pat. No. 6,332,088, and C. J. G. Bakker, R. M. Hoogeveen, J. Weber, J.J. van Vaals, M. A. Viergever, W. P. Th. M. Mali, “MR-guidedendovascular interventions: susceptibility-based catheter and nearreal-time scan technique”, Radiology 202, 273-276, 1996), and embeddingtuned RF coils in the catheter tip (eg: U.S. Pat. No. 6,289,233).

As previously mentioned, MRI is often the preferred method formonitoring the progress of surgery compared to X-ray and ultrasound.There are several reasons for this, including: no ionising radiation (sotheatre staff do not have to wear protective heavy lead clothing, whichis particularly important for intricate brain or cardiac operationswhich may last several hours); MRI generates an undistorted 3D image(rather than a projection, with no depth information in the case ofX-ray); MRI allows far better soft tissue characterisation anddifferentiation; MRI methods exist to monitor changes in tissueintegrity, based on diffusion, perfusion and/or flow: for example MRItechniques exist to monitor temperature, (which is particularly usefulduring thermal tissue ablation or cryosurgery), cell populations andcell chemistry; subsequent MR images directly show changes caused byinflammation, internal bleeding, thrombosis, organ motion or the directresults of surgery, etc.; MRI contrast agents can also be used tohighlight tissue changes (for example, this can be used to confirm thatall of a malignant tumour has been removed before completing theoperation). In specialised cases useful spectroscopic information canalso be obtained from the MR image.

Significant efforts have been made to combine various benefits affordedby these technologies. For example the steering system of U.S. Pat. No.6,241,671 incorporates X-ray fluoroscopy hardware (modified for use instrong magnetic fields) to monitor the catheter position in real-time.The Philips XMR system features a largely conventional MRI solenoid fortissue imaging and an X-ray fluoroscopy system for catheter positionmonitoring: the patient is placed on a moveable table that can be slidon rails between the two systems. In this way spatial registration ofthe MR and X-ray images is maintained. The latter system has theadvantage over the first of providing high quality MR images whenrequired, but it does not provide a catheter steering facility.Furthermore, it is rather cumbersome and clearly less than ideal to movethe patient between the two systems when carrying out a delicate lengthyoperation with the patient connected to life support and monitoringsystems.

Both of these combined systems use separate hardware for the tasks ofimaging the tissues, imaging the catheter and steering the catheter.

WO-A-99/18852 discloses the use of an MRI system for steering acatheter. The MRI system comprises a standard solenoid within which aworking volume is defined while the catheter is provided with a set oforthogonal coils which can be selectively activated by the user so as tointeract with the magnetic field generated by the solenoid to cause theorientation of the tip of the catheter to change. In some cases, thecatheter tip is provided with a RF receiver and transmitter.

Although this system overcomes some of the problems of the prior artmentioned above it still requires a relatively complex steeringapparatus to be provided on the catheter tip, and is furthermore notwell suited to allow easy access of a surgeon to a patient locatedwithin the working region of a conventional closed MRI magnet, being asolenoid or C-magnet.

In accordance with the present invention, a catheter imaging andsteering assembly comprises a magnetic field generating assemblyoperable in a first mode to generate a first magnetic field in a workingvolume located outside the assembly, the first magnetic field beingsuitable for use in a catheter steering procedure, and in a second modeto generate a second, static magnetic field in the working volumesuitable for conducting a magnetic resonance imaging process (MRI), thesecond magnetic field being weaker but more uniform in the workingvolume than the first magnetic field; and a catheter having a magneticseed attached whose orientation, and hence the steering direction of thecatheter, is determined by interaction with the first magnetic field.

We have devised a new form of imaging and catheter steering assemblywhich has the advantage of utilizing the same magnetic field generatingassembly for both imaging and steering, but has the added advantages ofutilizing a working volume located outside the assembly and a simpler“magnetic seed” attached to the catheter. This latter facility reducesthe complexity of hardware which needs to be provided at the cathetertip.

Consequently, the invention provides an assembly that combines thefacility for MR imaging of tissues when required (e.g.: to evaluate theresults of surgical intervention), with the ability to locate thecatheter (preferably in the same images, thus obviating the need forlater fusion of images from two different pieces of hardware, with thepossibility for spatial mis-registration), while also allowing remotemagnetic guidance of the catheter when required within the imagingvolume. In addition, it would be most preferred that the apparatusallowed surgeons free and uninhibited access to the patient, and thatthe patient should not have to be moved during surgery.

In the steering mode the assembly provides a “vector rotate magneticfield” which is used to steer a catheter tip equipped with a magneticseed. A “vector rotate magnetic field” is the term used for a magneticfield projected by a system of three (or more) electromagnets to cover aremote working region where steering is to take place. Typically theregion is roughly spherical and sufficiently large to cover asignificant proportion of the patient's body (e.g.: a 40 cm diametersphere). Within this region the field has sufficient flux density toinduce enough torque in a suitably sized “magnetic seed” to cause it torotate within the body, overcoming the friction and obstruction forcesfrom nearby body fluids and tissue structures, (about 0.2 T issufficient, as will be explained later). The direction of the field canbe selected over 4π solid radians by adjusting the currents in the threeelectromagnets, each of which generates a field component along oneorthogonal axis (see U.S. Pat. No. 6,241,671). In all prior art steeringsystems the magnetic field within the imaging volume is unsuitable forimaging, being far too inhomogeneous.

The “magnetic seed” can be a simple, passive magnetic element or anactive element whose magnetization can be locally controlled.

Examples of apparatus relevant to this invention can be found in WO02/49705, WO 02/43797, and WO 02/56047.

An example of an imaging and catheter steering assembly according to theinvention will now be described with reference to the accompanyingdrawings, in which:—

FIG. 1 is a schematic perspective view of a one-sided imaging magnet;

FIG. 2 is a schematic cross-section of the magnet, showing positions ofthe coils;

FIG. 3 is an example of suitable magnet circuit;

FIG. 4 is a perspective view of X axis gradient coil for imaging in theplane of the magnet;

FIG. 5 is a perspective view of Y axis gradient coil for imaging in theplane of the magnet;

FIG. 6 shows the position of the X axis gradient coil relative to mainmagnet in perspective view;

FIG. 7 is a perspective view of a modified X axis gradient and steeringcoil of the preferred embodiment;

FIG. 8 shows dimensions of a modified Y axis gradient and steering coilof the preferred embodiment;

FIG. 9 is a cross-sectional view showing positions of main magnet coilsand modified gradient/steering coils;

FIG. 10 is an example of a suitable circuit for combined X, Y steeringand gradient coils;

FIG. 11 is a schematic representation of a catheter steering system;and,

FIG. 12 is an illustration of the catheter steering assembly in moredetail.

The present invention describes an arrangement of electromagnets,(preferably wound from high-temperature superconducting wire, optimisedfor high rate of change of current, ie: “high dI/dt” or “AC capablewire” and placed within a suitable cryostat (not shown) in which thecurrents can be adjusted under user control to allow MR images to beacquired from a remote sample volume in a second mode of operation, andto orient a magnetic seed embedded in a catheter tip withinsubstantially the same volume for the purpose of steering the catheterin a first mode.

In the following description a preferred embodiment is described. It isnot intended to limit the application to this embodiment, which isamenable to scaling and other geometry changes.

FIG. 1 shows an arrangement of co-planar electromagnet coils (about 3 min diameter) for generating a 14 cm diameter uniform spherical workingvolume (DSV, 7) with flux density 0.1 T, uniform to 50 ppm. The DSV 7 isdisplaced by 21 cm to one side of the coil. MR imaging is possiblewithin the DSV, using suitable gradient and RF fields, as will bedescribed shortly. The Cartesian co-ordinate system (7 a) is usedthroughout the description, and has its origin at the centre of the DSV.The coil positions, dimensions and currents required for imagingread-out mode are given in Table 1, which should be read with referenceto FIG. 2 (showing the coil positions in cross-section). TABLE 1 Flat0.1T magnet for one-sided imaging system (current directions shown forread-out mode). Current Coil # a1 (cm) a2 (cm) b1 (cm) b2 (cm) Turns (A)1 121.18 125.28 27.18 52.82 1051 100 2 125.28 146.82 27.18 52.82 5523100 3 76.24 93.77 21.24 38.77 3073 −100 4 38.86 41.14 28.86 31.14 520100 5 29.08 30.92 29.08 30.92 339 −100 6 19.64 20.36 29.64 30.36 52 100

The flat electromagnet is composed of two distinct coil sub-circuits (orsub-magnets), one set comprising coils 1, 3 and 5, and one comprisingcoils 2, 4, and 6. (FIG. 3). Thus, as shown in FIG. 3, superconductingcoils 2, 4, and 6 are connected in series to form the constant-fieldsub-magnet. After energisation by a DC power supply 16, asuperconducting switch 15 is closed and the power supply 16 may beremoved. Current continues to flow in the constant field sub-magnet. Asecond DC power supply 13 is connected to the variable-field sub-magnet(coils 1, 3, and 5) by an H-bridge inverter (switch pairs 11 and 12).These switches may be solid state devices, such as IGBTs or relays. Whenswitch pair 11 are closed current flows in the variable-fieldsub-magnet, generating the pre-polarization field. The current in theconstant field sub-magnet is unaffected due to the low coupling(coupling coefficient K^(˜)1%) between the magnets. After allowing asufficient duration for the magnetization of the sample to build up,switches 11 are opened. The energy stored by the coils 1, 3 and 5 keepsthe current flowing and charges a capacitor 14. The value of thiscapacitor is chosen so that it resonates with the self inductance of thevariable field sub-magnet at the frequency defined by 1/TS, where TS isthe desired field-switching time. For example, in the preferred 0.1 Tmagnet, the inductance of the variable-field sub-magnet is 21.5 Henries.For a switching time of 100 ms the capacitor 214 needs to be 47 uF. Whenthe current has fallen to zero (the stored energy has exchanged from themagnet to the capacitor), the switches 12 are closed. Current thenincreases through the variable field sub-magnet in the oppositedirection until it reaches the same level as before, but flowing in theopposite direction. Most of the energy is supplied from the capacitor,with the power supply making up any small losses. Switch 17 is thenopened (the current flowing through it having fallen to zero). Afterapplying the imaging pulse sequence, the current is again reversed byclosing switch 17, opening switch pair 12, waiting until the currentthrough the coils reaches zero, closing switch pair 11, waiting untilthe current through the coils reaches the peak value, then openingswitch 17. This regenerative switching process may be repeated as oftenas required. Power supply 13 only needs to supply a small quantity ofpower to make up for any losses.

It is not a strict requirement to reverse the current to the sameabsolute level when flowing in the opposite direction—the correctioncoils could simply be turned off, or the current simply reduced. Theformer option would reduce the stored energy to 0.7 MJ during thepre-magnetization period, but there would only a 2.7 fold enhancement infield and NMR signal. However, there are good technical reasons forpreferring to have the current rise to the same absolute level butflowing in the opposite sense in the pre-polarization period compared tothe read-out period: the stored energy in both pre-polarization andread-out modes with reversed but equal-magnitude current is the same, sochoosing any lower current value means that energy will have to beremoved, stored for the duration of the imaging sequence, which could bea few seconds, then added back into the magnet. This is inconvenient:0.1 MJ of energy would have to be stored in the example above. Resonantswitching of the current using a capacitor, as described above, is themost convenient way of rapidly reversing the current to the same value,and temporarily storing the energy in the capacitor bank during theshort change-over period TS. This is described in more detail in WO02/56047.

The coil positions and turns-densities are arranged so that there issubstantially zero flux linkage between the two sub-magnets. This meansthat the sub-magnets have substantially zero mutual inductance, anecessary requirement if they are to be treated as electrically isolatedfrom the perspective of electronics design. However, with increasedcomplexity in the electronics it is conceivable that sub-magnets withnon-zero coupling could be used.

The present invention uses the same hardware to provide the steeringvector rotate field for catheter steering. In particular, the mainmagnet provides the variable Z-component, and the X and Y gradient coilscan provide the X and Y components of the vector rotate field.

The steering ability arises from the magnetic seed experiencing a torquetrying to align it with the applied field. In steering mode gradients inthe magnetic field within the DSV also undesirably impose translationalforces on the seed, but these are insignificant compared to the twistingtorque.

The torque on a small permanent magnet in a catheter tip having magneticmoment m in a magnetic field with flux density B is:

-   -   Γ=qm×B

The magnetic moment m of a cylindrical seed of permanent magneticmaterial with remanence B_(seed), radius r and length l is:$m = {l\quad\pi\quad r^{2}\frac{B_{seed}}{\mu_{0}}}$

Experiment shows that the torque needed to deflect the tip is typicallyof the order 10 gramme centimetres (maximum torque occurs when the seedmagnetisation is orthogonal to the applied magnetic field). Assumingtypical values for the magnetic seed of r=1 mm, l=5 mm and B_(seed)=0.4T (typical for hard ferrite), the required steering flux density isfound to be 0.2 T. A hard ferrite seed can be de-magnetised by applyingan oscillating decaying current to coil wound around the seed. The seedmay be re-magnetised by a current pulse. This feature is described morefully below. It will be advantageous to use the ability to turn off themagnetic seed during imaging mode, and re-activate it for steering mode.

To generate the Z-component of the steering magnetic field vector fromthe main magnet coils with magnitude ranging from −0.2 T to +0.2 T it isnecessary to adjust the coil currents to intermediate values compared tothe currents needed in imaging mode. In comparison to imaging mode, theintermediate current values needed in steering mode require storedenergy to be removed or added to the electromagnets using the powersupplies. For example, the Z-component of steering field can be adjustedbetween −0.2 T to +0.2 T by setting the current in coils 2, 4 & 6 at theappropriate value between +74 amps and −74 amps, and turning off thecurrent in coils 1, 3 & 5.

The maximum field gradient generated by the main magnets in steeringmode occurs at either extreme of the −0.2 to +0.2 T range, and is 45G/cm. The maximum force experienced by a magnetic seed of moment m in amagnetic field gradient is given by:

-   -   F=−m.grad(B)

For the seed in this example placed in the 45 G/cm gradient the maximumtranslational force is therefore 2.25 mN, which is negligible.

The X and Y components of the steering field are generated by flat X andY gradient coils whose principles will be described with reference toFIGS. 4 to 6. These “kidney-shaped” gradient coils 17,20 lie in theplane of the magnet and generate linear dBz/dx and dBz/dy fields acrossan XY plane slice through the centre of the DSV. The fields generated bythese coils also contain orthogonal field components, in the X and Ydirections respectively. These orthogonal components have negligibleeffect on the MR image. FIG. 4 shows the X gradient coils (17), and FIG.5 the Y gradient coils (20). The field orientation at the centre of theDSV is shown by the vectors (18 and 21). Also shown in FIGS. 4 and 5 arecontour plots of the variation in the Z-component of the field generatedby the coils across a 20×20 cm slice through the DSV (19 & 22),demonstrating the linearity of the gradient fields. The pair of kidneycoils in each set are connected in series (not shown), with currentorientation as shown by the arrows. FIG. 6 shows the X gradient coils,17, in position relative to the main magnet coils, 1 to 6. The Xgradient set is displaced by 360 mm along the Z axis and lies justbehind the inner magnet coils (numbers 4, 5 & 6), as shown in FIG. 6.The Y gradient set (not shown) sits just behind the X gradient set.

To generate 0.2 T X and Y steering field components from these coilsrequires very much higher amp-turn values compared to operation asgradient coils. For example, to generate 0.2 T field requires about halfa million amp-turns, in comparison to one thousand amp-turns needed togenerate imaging gradient fields. Current densities this high cannot beachieved in reality with coils of the thin cross section shown in FIGS.4 to 6. It is therefore necessary to increase the cross-sectional areaof the coils. FIG. 7 shows a close up perspective view of the new“fatter” combined X-gradient and steering coils (24), and a contour plotof the X field component across the same 20×20 cm slice through the DSV(23). The variation in magnitude of the X component across the slice isonly 15%, so the magnetic seed will experience only this variation intorque as its position within the slice varies. FIG. 8 shows thedimensions of the combined gradient and steering coil in the preferredembodiment. FIG. 9 is a cross-sectional view in the YZ plane showing therelative positions of the X and Y gradient/steering coils (24, 25) andthe inner magnet coils (4, 5 & 6). The current density in each of thenew gradient coils is about 200 amps/mm² and the winding cross sectionsare 50×50 mm for the X coil and 50×80 mm for the Y coil. The Y coilrequires more amp-turns to generate the 0.2 T field because it isfurther from the DSV; it is therefore deeper than the X coil in the Zdirection.

In practice to achieve the desired amp-turns without overheating, thesecoils must be made from superconducting wire, preferably hightemperature superconductor (HTS). Second generation YBaCuO HTS wire istypically available as tape and is most easily wound into “pancakecoils” (in which tape is wound upon itself to create a coil which is fatin the winding direction but only one tape-width thick). Assuming tapedimensions of 10 mm wide by 0.5 mm thick (including insulation) a 100turn pancake coil would be 50 mm thick in winding direction and 10 mmwide. Therefore the X coil would require a stack of 5 pancakes connectedin series and the Y coil a stack of 8. The total turns of the X coil istherefore 5×100=500 implying a coil current of 1000 amps is needed toachieve a 0.2 T field. This is achievable with a next generation HTSconductor.

FIG. 10 shows a possible circuit diagram for the X and Ygradient/steering coils. The physical layout of the coils results innear zero mutual inductance, so they can be treated aselectromagnetically isolated units. The inductances 24 and 25 representthe total series inductance of the X and Y steering/gradient coils.These are connected to power supplies 29 and 31 which supply thenecessary steady state current for steering mode. As previouslydescribed the coils are made from stacks of series connected pancakewindings. The pancake windings are preferably connected together so thatall coils lying in a plane are connected in series (ensuring currentflows in the correct sense, as shown in FIGS. 4 & 5), then connected inseries to the next layer. The coils are tapped (26 and 27) after theinnermost pancake windings of the X and Y coils that face each other atz=360 mm. The taps are connected to independent power supplies 28 and 30which supply the necessary fast rise time pulsed currents for gradientoperation. In this way only the innermost pancakes of the stack nearestto z=360 mm are used for gradients. This offers best performance (thetapped coils have lower inductance coils allowing faster gradient pulserise times, and the thinner cross section results in better gradientfield linearity).

Clearly the superconducting gradient/steering coils will need to becooled below their critical temperature. This can be achieved by placingthem inside the magnet cryostat. The radio-frequency (RF) coils used fortransmitting RF pulses and receiving NMR signals may be placed withinthe same cryostat, preferably in front of the inner magnet coils, closeto the DSV. In this case the cryostat will require an RF transparentwindow. This is also described in more detail in WO 02/56047.

Any conventional imaging pulse sequence can be used.

As already explained, in addition to imaging, the assembly is used forsteering a catheter. An example of a suitable catheter is shown in FIGS.11 and 12. FIG. 11 shows a catheter steering assembly generallyindicated at 61, the catheter steering assembly comprising a housing 62within which is enclosed a sphere 63 of hard or semi-hard magneticmaterial such as ferrite, the sphere being enclosed by three orthogonalelectric microcoils 64 a, 64 b, 64 c.

A guide wire 65 is connected to the housing 62, the guide wirecontaining electrical lines 66 a, 66 b, 66 c for supplying electricsignals to the electrical coils 64 a, 64 b, 64 c respectively. The guidewire 65 passes through a catheter generally indicated at 67, thecatheter having an elongate body and a central bore 68 through which theguide wire passes.

At the end of the catheter body closest to the catheter steeringassembly 61, an annular lip 69 is provided so as to narrow the diameterof the bore 68 to form an opening 70. At a predetermined distance alongthe guide wire from the catheter steering assembly 61, a disk 71 isattached to the guide wire 65, the radius of the disk being arranged tobe just less than that of the internal diameter of the catheter 67 andyet larger than the diameter of the opening 70. During use, theattachment of the disk to the guide wire 65, prevents the cathetersteering assembly 61 from separating from the catheter 67 by more than apredetermined distance. This distance can be arranged according to theuse of the catheter in question.

The catheter guide assembly 61 and catheter 67, along with the guidewire 65, are formed from suitable materials to be used within the bodyof a living subject such as the human body. The guide wire 65 is leadout of the body and is adapted for manipulation by a surgeon. In thisexample the guide wire has sufficient stiffness to allow the cathetersteering assembly 61 and catheter 67 to be moved through body cavitiesor lumens by applying a sufficient axial force to the guide wire 65.

The electrical lines 66 a, 66 b, 66 c are attached to an external signalgenerator 75 which is adapted to provide electrical signals to therespective electrical lines 66 a, 66 b, 66 c. The signal generator 75 iscontrolled by a computer 76 having a processor operating controlsoftware. An appropriate input device 77 such as a keyboard or joystickallows the surgeon to control the electrical signals being passed to thecatheter steering assembly 61 using the computer 76.

The magnet of FIG. 1 is schematically represented at 78 and ispositioned so as to apply a magnetic field with which the cathetersteering assembly 61 may interact.

FIG. 12 shows the catheter steering assembly 61 in more detail, with thehousing 62 removed. The ferrite sphere 63 is encircled by the threeelectric coils 64 a, 64 b, 64 c. Each of these coils comprises a numberof turns of high conductivity electrical wire, the coils beingelectrically connected to the electrical signal generator 75 using thecorresponding electrical lines 66 a, 66 b, 66 c positioned along theguide wire 65.

As indicated in FIG. 12, the three coils are arranged about the centreof the sphere 63 along mutually orthogonal axes. If sufficientlyisotropic ferrite is used for the sphere 63, then the arrangement of thecoils 64 a, 64 b, 64 c in this manner allows the generation of amagnetic field within the ferrite in an arbitrary direction bysuperposition of the fields generated by each coil individually. Thismay be achieved by applying one or more suitable current pulses to oneor more of the coils such that the combined magnetic field generated bythe current in the coils is greater than the coercive force required tomove the magnetic domains within the material.

In this manner, not only can the direction of the magnetisation bechanged at will, the magnitude and polarity of this magnetisation canalso be controlled. An additional benefit is that a decaying oscillatingpulse of current applied to one or more of the microcoils candemagnetize the ferrite for subsequent MRI imaging.

1. An imaging and catheter steering assembly comprising: a magneticfield generating assembly operable in a first mode to generate a firstmagnetic field in a working volume located outside the assembly, thefirst magnetic field being suitable for use in a catheter steeringprocedure, and in a second mode to generate a second, static magneticfield in the working volume suitable for conducting a magnetic resonanceimaging process (MRI), the second magnetic field being more uniform inthe working volume than the first magnetic field; and a catheter havinga magnetic seed attached whose orientation, and hence the steeringdirection of the catheter, is determined by interaction with the firstmagnetic field.
 2. An assembly according to claim 1, wherein themagnetic field generating assembly comprises first and secondelectromagnets whose currents may be adjusted or reversed within aworking range including zero, so as to vary the magnitude of the Zdirection component of a vector rotate magnetic field in the workingregion as required in the steering procedure of the first mode, and mayalso be adjusted to a fixed setting which, in combination with the firstand second electromagnet's coil positions and turns densities, generatesa relatively uniform magnetic field in the working region as required inthe imaging procedure of the second mode.
 3. An assembly according toclaim 2, further comprising third and fourth magnets for generatingpulsed magnetic fields with linear gradients in mutually orthogonal Xand Y directions respectively, orthogonal to the Z direction, during theMR process and substantially static magnetic fields in the X and Ydirections respectively during the steering procedure.
 4. An assemblyaccording to claim 2, wherein one or more of the magnets aresuperconducting electromagnets.
 5. An assembly according to claim 4,wherein the coils are made from high temperature superconductor.
 6. Anassembly according to claim 2, wherein the first and second magnetscomprise electrical coils, the coil positions and turns densities beingsuch that the coils exhibit substantially zero mutual inductance.
 7. Anassembly according to claim 1, wherein the magnetic dipole of themagnetic seed in the catheter tip may be switched on in steering modeand off in imaging mode.
 8. An assembly according to claim 7, whereinthe magnetic seed comprises a semi-hard permanent magnet whosemagnetisation may be switched on by applying a pulse of current to amicrocoil wound around the seed, and switched off by applying a decayingoscillating pulse of current to the same microcoil.
 9. A method ofcarrying out a medical procedure, the method comprising: providing animaging and catheter steering assembly comprising a magnetic fieldgenerating assembly operable in a first mode to generate a firstmagnetic field in a working volume located outside the assembly, thefirst magnetic field being suitable for use in a catheter steeringprocedure, and in a second mode to generate a second, static magneticfield in the working volume suitable for conducting a magnetic resonanceimaging process (MRI), the second magnetic field being more uniform inthe working volume than the first magnetic field; and a catheter havinga magnetic seed attached whose orientation, and hence the steeringdirection of the catheter, is determined by interaction with the firstmagnetic field; inserting the catheter into a body; steering thecatheter through the body by selectively operating the assembly in thefirst mode; and obtaining an image of part of the body by operating theassembly in the second, imaging mode.